This invention relates to coherent imaging using vibratory energy, such as ultrasound and the like and, in particular, to ultrasound imaging of flowing tissues such as blood.
There are a number of modes in which ultrasound can be used to produce images of objects. The ultrasound transmitter may be placed on one side of the object and the sound transmitted through the object to the ultrasound receiver placed on the other side ("transmission mode"). With transmission mode methods, an image may be produced in which the brightness of each pixel is a function of the amplitude of the ultrasound that reaches the receiver ("attenuation" mode), or the brightness of each pixel is a function of the time required for the sound to reach the receiver ("time-of-flight" or "speed of sound" mode). In the alternative, the receiver may be positioned on the same side of the object as the transmitter and an image may be produced in which the brightness of each pixel is a function of the amplitude or time-of-flight of the ultrasound reflected from the object back to the receiver ("refraction", "backscatter" or "echo" mode). The present invention relates to a backscatter method for producing ultrasound images.
There are a number of well known backscatter methods for acquiring ultrasound data. In the so-called "A-scan" method, an ultrasound pulse is directed into the object by the transducer and the amplitude of the reflected sound is recorded over a period of time. The amplitude of the echo signal is proportional to the scattering strength of the refractors in the object and the time delay is proportional to the range of the refractors from the transducer. In the so-called "B-scan" method, the transducer transmits a series of ultrasonic pulses as it is scanned across the object along a single axis of motion. The resulting echo signals are recorded as with the A-scan method and either their amplitude or time delay is used to modulate the brightness of pixels on a display. With the B-scan method, enough data are acquired from which an image of the refractors can be reconstructed.
In the so-called C-scan method, the transducer is scanned across a plane above the object and only the echoes reflecting from the focal depth of the transducer are recorded. The sweep of the electron beam of a CRT display is synchronized to the scanning of the transducer so that the x and y coordinates of the transducer correspond to the x and y coordinates of the image.
Ultrasonic transducers for medical applications are constructed from one or more piezoelectric elements sandwiched between a pair of electrodes. Such piezoelectric elements are typically constructed of lead zirconate titanate (PZT), polyvinylidene difluoride (PVDF), or PZT ceramic/polymer composite. The electrodes are connected to a voltage source, and when a voltage is applied, the piezoelectric elements change in size at a frequency corresponding to that of the applied voltage. When a voltage pulse is applied, the piezoelectric element emits an ultrasonic wave into the media to which it is coupled at the frequencies contained in the excitation pulse. Conversely, when an ultrasonic wave strikes the piezoelectric element, the element produces a corresponding voltage across its electrodes. Typically, the front of the element is covered with an acoustic matching layer that improves the coupling with the media in which the ultrasonic waves propagate. In addition, a backing material is disposed to the rear of the piezoelectric element to absorb ultrasonic waves that emerge from the back side of the element so that they do not interfere. A number of such ultrasonic transducer constructions are disclosed in U.S. Pat. Nos. 4,217,684, 4,425,525, 4,441,503, 4,470,305 and 4,569,231, all of which are assigned to the instant assignee.
When used for ultrasound imaging, the transducer typically has a number of piezoelectric elements arranged in an array and driven with separate voltages (apodizing). By controlling the time delays (or phase) and amplitude of the applied voltages, the ultrasonic waves produced by the piezoelectric elements (transmission mode) combine to produce a net ultrasonic wave focused at a selected point. By controlling the time delays and amplitude of the applied voltages, this focal point can be moved in a plane to scan the subject.
The same principles apply when the transducer is employed to receive the reflected sound (receiver mode). That is, the voltages produced at the transducer elements in the array are summed together such that the net signal is indicative of the sound reflected from a single focal point in the subject. As with the transmission mode, this focused reception of the ultrasonic energy is achieved by imparting separate time delays (and/or phase shifts) and gains to the signal from each transducer array element.
This form of ultrasonic imaging is referred to as "phased array sector scanning", or "PASS". Such a scan is comprised of a series of measurements in which the focused ultrasonic wave is transmitted, the system switches to receive mode after a short time interval, and the reflected ultrasonic wave is received and stored. Typically, the transmission and reception are steered in the same direction (.theta.) during each measurement to acquire data from a series of points along a scan line. The receiver is dynamically focused at a succession of ranges (R) along the scan line as the reflected ultrasonic waves are received. The time required to conduct the entire scan is a function of the time required to make each measurement and the number of measurements required to cover the entire region of interest at the desired resolution and signal-to-noise ratio. For example, a total of 128 scan lines may be acquired over a 90 degree sector, with each scan line being steered in increments of 0.70.degree.. A number of such ultrasonic imaging systems are disclosed in U.S. Pat. Nos. 4,155,258, 4,155,260, 4,154,113, 4,155,259, 4,180,790, 4,470,303, 4,662,223, 4,669,314 and 4,809,184, all of which are assigned to the instant assignee.
The use of ultrasonic imaging systems to measure the velocity of moving tissues such as blood has many clinical applications. For example, an image of the heart in which pixel brightness is determined by the velocity of the reflecting tissues can provide valuable information concerning the mechanical motion of the heart and the flow of blood through its chambers during a cardiac cycle. There are two well known methods for producing such images using phased array sector scanning techniques--Doppler measurements and correlation measurements.
Ultrasonic echoes backscattered from moving tissues such as blood are frequency shifted by an amount proportional to the frequency of the ultrasonic sound and the velocity of the blood flow. Movement of red blood cells through the sample volume toward the transducer array compresses the wavelength of the reflected echo, increasing its frequency, whereas movement of red blood cells away from the transducer array lengthens the wavelength of the echo, decreasing its frequency. The Doppler method measures only the component of mean velocity in the same direction as the transmitted ultrasonic beam. The formula relating Doppler frequency shift to velocity in the beam direction is: EQU .DELTA.f=2f.sub.0 v/c
where
.DELTA.f=frequency shift, PA0 f.sub.0 =ultrasonic carrier frequency, PA0 v=mean velocity of flow in the beam direction, PA0 c=speed of sound in the tissue (1450 mm/sec).
In typical systems using a carrier frequency of 2-5 MHz, human blood flows produce Doppler frequency shifts of about 0.2 to 8 kHz. Ultrasonic imaging systems which produce Doppler flow images are described in U.S. Pat. Nos. 4,217,909; 4,265,126; 4,182,173; 4,257,278; and 4,530,363, all of which are assigned to the instant assignee.
The second method used to produce ultrasonic flow images correlates the signals produced by two echoes to determine the distance moved by the blood cells between the two ultrasonic beam transmissions. The two transmissions may be directed in the same beam angle and the amplitude of the sampled echo signals correlated with each other to determine the amount of shift in position which takes place between transmissions along the beam direction. Since the time interval between transmissions is known, this measured displacement can be used to calculate velocity. The same method can be used to measure flow velocity across the direction of the beam by emitting a series of beams at different beam angles and then correlating the amplitudes of the received echo signals in both the beam direction and the cross beam direction. Such correlation ultrasonic imaging systems are described in U.S. Pat. Nos. 4,587,973 and 4,567,898, which are assigned to the instant assignee, and in articles by O. Bonnefous et al. in Ultrasonic Imaging 8, pp. 73-85 (1986) entitled "Time Domain Formulation Of Pulse-Doppler Ultrasound And Blood Velocity Estimation By Gross Correlation"; by P. M. Embree et al. in IEEE 1985 Ultrasonics Symposium, pp. 963-66 entitled "The Accurate Ultrasonic Measurement Of The Volume Flow Of Blood By Time Domain Correlation"; and by Gregg E. Trahey et al. in IEEE 1987 Ultrasonic Symposium, pp. 957-61, entitled "Measurement of Local Speckle Pattern Displacement to Track Blood Flow in Two Dimensions." To correlate in the cross beam direction, however, a plurality of beams at different beam angles must be produced to provide a magnitude envelope in the cross beam direction and the process must be repeated again to produce a second envelope which can be correlated with the first envelope.